JP2014161590A - Dental x-ray imaging apparatus and image correction method in dental x-ray imaging - Google Patents

Abstract

PROBLEM TO BE SOLVED: To reliably remove or reduce a faulty shadow included in a panoramic image of an imaging target by correction processing in an imaging system which comprises an X-ray tube and a detector disposed across the subject and is rotational around the subject.SOLUTION: An X-ray imaging apparatus comprises an X-ray tube and a detector. The X-ray tube and the detector are rotated with a jaw part of a subject positioned therebetween for scanning by X rays (steps S1, S2). A first panoramic image with one of two structures, e.g. a tooth row, in focus is captured (step S2). A second panoramic image with the other of the two structures, e.g. the cervical spine, in focus is captured (step S3). With the one structure being an imaging target, an image of the other structure (faulty shadow) included in the panoramic image of the imaging target is removed or reduced on the basis of the first and second panoramic images, to generate a panoramic image of the imaging target (steps S4 to S7).

Description

  The present invention is a dental X-ray that reconstructs a panoramic image by scanning a jaw of a subject with an X-ray beam to detect transmitted X-rays and processing the detected X-ray data by a tomosynthesis method. The present invention relates to a line imaging apparatus and an image correction method in dental X-ray imaging.

  In recent years, tomography of a subject by the tomosynthesis method has been actively performed. Although the principle of this tomosynthesis method has been known for a long time (see, for example, Patent Document 1), in recent years, a tomographic method has also been proposed in which it is desired to enjoy the simplicity of image reconstruction based on the tomosynthesis method. (For example, see Patent Document 2 and Patent Document 3). In addition, many examples are seen for dental use (see, for example, Patent Document 4 and Patent Document 5).

  As one application of the tomosynthesis method to dentistry, a panoramic imaging apparatus that obtains a panoramic image in which a curved dentition is developed in a two-dimensional plane is generally put into practical use. This panoramic imaging apparatus normally draws a pair of an X-ray tube and a detector having pixels over a longitudinal width around the oral cavity of a subject to form a constant trajectory along a dentition whose rotation center is assumed. Thus, a mechanism for rotating the rotation center in a complicated manner is provided. A constant value is maintained between the X-ray tube and the detector. The above-described constant trajectory is a trajectory for focusing on a reference tomographic plane (a tomographic plane existing three-dimensionally) set in advance along a dentition regarded as a standard shape and size. During this rotation, X-rays emitted from the X-ray tube are transmitted through the subject at regular intervals and detected as digital frame data by the detector. For this reason, frame data focused on the reference tomographic plane is collected at regular intervals. This frame data is reconstructed by the tomosynthesis method to obtain a panoramic image of the reference tomographic plane.

  Patent Document 6 discloses an example of a panoramic imaging apparatus having an imaging system in which an X-ray tube and a detector can both rotate independently of each other so as to draw a circular orbit around the same center point. . The jaw is positioned in the circular orbit. X-rays irradiated from the X-ray tube are controlled so as to always face the detection surface of the detector.

JP-A-57-203430 JP-A-6-88790 JP-A-10-295680 US Patent Publication US2006 / 0203959 A1 JP2007-136163A International Publication WO2012 / 008492

  In the case of the panoramic imaging apparatus described in Patent Document 6 described above, the distance between the X-ray tube and the detector is not fixed, and the distance changes for each X-ray path irradiated by scanning. This makes it possible to set a desired X-ray path that avoids the cervical vertebra and the left and right upper jaws as much as possible, while making the control of the angular velocity on the circular orbit that rotates the X-ray tube and detector relatively complicated. Become.

  On the other hand, it is possible to easily design an angular velocity that realizes an X-ray path that avoids structures such as the cervical vertebrae that are obstructive to the imaging of the dentition, but the reconstruction of the desired tomographic plane by the tomosynthesis method, Shift and add processing is complicated. Moreover, in order to perform this reconstruction process more accurately, various calibrations are performed using a phantom in order to grasp the relationship between the distance in the tomographic direction and the amount of shift-and-add, that is, the structure of the imaging space. There is a need. That is, the calibration process is complicated and takes time. In addition, only the collection scenes corresponding to the number of calibrations can be dealt with, and the X-ray tube and the detector can be freely rotated independently of each other, that is, the distance between the X-ray tube and the detector is It can be said that it is cumbersome to effectively use the advantage of the folding angle of the variable imaging system.

  Even if we dare to endure such inconvenience, it is impossible to design the trajectory so that the X-ray beam does not completely pass through the cervical vertebra, which becomes an obstacle shadow. In addition, when it is desired to observe a panoramic image of a cross section passing through the cervical spine, the dentition is reflected in the panoramic image of the cervical spine as an obstacle shadow. Furthermore, the site that becomes the shadow of the obstacle is not limited to the dentition and the cervical spine, but may be the left and right jawbones.

  Under such circumstances, there has been a demand for a method for removing or reducing the obstacle shadow that is inevitably reflected in the panoramic image of the target portion.

  The present invention has been made in view of the above circumstances. When imaging is performed by rotating the X-ray tube and the detector around the subject, the obstruction shadow that appears in the panoramic image of the object is reliably corrected. It is an object of the present invention to provide a dental X-ray imaging apparatus and an image correction method in dental X-ray imaging that can be further removed or reduced to further enhance the rendering ability of the object.

  In order to achieve the above object, an X-ray imaging apparatus according to an aspect of the present invention includes an imaging system including an X-ray tube that irradiates X-rays and an X-ray detector that detects the X-rays. This is a dental X-ray imaging apparatus that is rotated with the jaw of the subject positioned between the X-ray tube and the detector. The apparatus includes a first image obtaining unit that obtains a first panoramic image focused on one of the two structures existing on the jaw, and one of the two structures. Second image acquisition means for acquiring a second panoramic image focused on the remaining one structure, the one structure as an object to be imaged, and the object of the remaining one structure An object for generating a panorama image of the object from which the reflection of the obstacle shadow is removed or reduced based on the first and second panorama images when the object is reflected in the panorama image. And an image generation means.

As a preferred example, the object image generating means includes:
Scaling means for adjusting the size of the panoramic image of the cervical vertebra (second panoramic image) to the size of the panoramic image of the dentition (first panoramic image), and the panoramic image of the cervical vertebra adjusted by the scaling means A blur image generating unit that generates a blurred image of the cervical vertebra by superimposing and integrating a blur function defined by the focal plane on each pixel; and a panoramic image of the dentition acquired by the first image acquiring unit; The pixel value is subtracted (in the case of a pixel value obtained by natural logarithm) or divided (natural logarithm) for each pixel with respect to the blurred image of the panoramic image of the cervical vertebra generated by the blurred image generation unit. Difference image calculation means for providing the difference image as a panoramic image of the object after the obstacle shadow is removed or reduced.

  An image correction method in dental X-ray imaging according to another aspect of the present invention includes an imaging system including an X-ray tube that emits X-rays and an X-ray detector that detects the X-rays. This is carried out in dental X-ray imaging in which the jaw of the subject is rotated between the X-ray tube and the detector. A first panoramic image focused on one of the two structures present in the jaws was acquired and the remaining one of the two structures was focused When the second panoramic image is acquired, the one structure is set as an imaging target, and the reflection of the remaining one structure on the panoramic image of the target is set as an obstacle shadow, the first And generating a panoramic image of the object from which the reflection of the obstacle shadow is removed or reduced based on the second panoramic image.

  According to the present invention, when imaging is performed by rotating the X-ray tube and the detector around the subject, the obstacle shadow that appears in the panoramic image of the object is reliably removed or reduced by the correction process. Thereby, the ability to depict the object can be further enhanced.

In the accompanying drawings,
FIG. 1 is a perspective view illustrating an overview of a panoramic imaging apparatus as an X-ray imaging apparatus according to an embodiment; FIG. 2 is a diagram for explaining an arrangement configuration of an X-ray tube and a detector of the panorama imaging apparatus, FIG. 3 is another diagram illustrating the arrangement configuration of the X-ray tube and the detector of the panoramic imaging apparatus together with the rotatable directions of the X-ray tube and the detector and their facing states. FIG. 4 is a diagram of an arrangement example illustrating the outline of the configuration of the detector, FIG. 5 is a block diagram illustrating the electrical configuration of the detector. FIG. 6 is a graph for explaining the relationship between an electrical pulse detected in response to the incidence of X-ray photons and a threshold value, which is given to a photon counting detector. FIG. 7 is a graph for explaining an example of the relationship between the energy spectrum with respect to the incidence frequency (count value) of X-ray photons and the energy region given by the discrimination circuit; FIG. 8 is a block diagram showing an outline of the electrical configuration of the entire panorama imaging apparatus, FIG. 9 is a flowchart schematically showing a flow of scan and obstacle shadow removal executed in the embodiment; FIG. 10 is a diagram showing a part of the process of removing the obstacle shadow, A graph showing an example of a gain curve for reconstructing a panoramic image of a dentition, A graph showing an example of a gain curve for reconstructing a panoramic image of the cervical spine, An image showing an example of a panoramic image of the dentition (before removing the shadow), An image showing an example of a panoramic image of the cervical spine, An image (an image of a patient different from the patient in FIG. 13) showing an example of the panoramic image of the dentition before removing the obstacle shadow, and It is an image which shows an example after the obstacle shadow removal (reduction) of the panoramic image of a dentition which should be compared with FIG.

  Embodiments of the present invention will be described below with reference to the accompanying drawings.

  An embodiment of a panoramic imaging device as an X-ray imaging device according to the present invention will be described with reference to FIGS.

  The panoramic imaging device 1 is configured as a dental device that captures a panoramic image of a subject's jaw (including a dentition). According to this apparatus, a pseudo three-dimensional cross-sectional image of the jaw of the subject (the image itself is a two-dimensional image, depending on the shape of the imaging region such as a dentition, etc., depending on the configuration and functions described below. A cross-sectional image displayed two-dimensionally) can be taken. Although it is configured as a panoramic imaging apparatus according to the present embodiment, it is not necessarily limited to the field of dentistry, and can be applied to various parts such as mammography, otolaryngology imaging, bones and joints of limbs. It can also be applied to uses such as corpse identification, home inspection, and nondestructive inspection for identification.

  FIG. 1 shows an appearance of a dental panoramic imaging apparatus 1 according to the present embodiment.

  The panoramic imaging apparatus 1 includes a pedestal 12 on which casters 11 are mounted, an elevating unit 13 and a power supply box 14 mounted on the pedestal 12, and a console 17. The elevating unit 13 includes an elevating mechanism (not shown) therein, and the upper elevating unit of the unit is configured to be movable up and down within a predetermined range electrically with respect to the base 12 (that is, the floor surface). Yes. If the vertical movement direction of the elevating unit 13 is the Z axis, XYZ orthogonal coordinates as shown in the figure can be assumed. The power supply unit 14 supplies necessary power to each part of the system.

  The panoramic imaging apparatus 1 also includes two arms 15 and 16 extending in the X-axis direction (that is, laterally) from the lifting unit of the lifting unit 13. The two arms 15 and 16 extend along the Y-axis direction. When viewed, both are formed in a substantially L shape, and one end of each of the arms 15 and 16 is superposed so as to overlap each other and attached to the side surface of the elevating unit. A rotation mechanism 13D is provided that can rotate the two arms 15 and 16 independently of each other, that is, at different speeds, and an X-ray tube 21 is provided at the distal end of each of the two arms 15 and 16. And a detector 22. A slit (diaphragm) 23 for forming X-rays into a fan shape is disposed on the front surface of the X-ray tube 21 on the X-ray irradiation side. The opening area is controlled by an opening drive unit 23D (see FIG. 8) such as a motor, which will be described later, and the X-ray tube 21 and the arms 15 and 16 by the rotation mechanism 13D and the arms 15 and 16. Support means for supporting the detector 22 so as to be driven independently of each other is configured.

  The X-ray tube 21 is configured as a rotary anode type X-ray tube used for an appropriate anode material such as tungsten. The X-ray tube 21 has a dotted X-ray focal point (for example, a diameter of 0.1 mm to 0.5 mm) FP. The X-ray tube 21 emits X-rays in response to driving power supplied from a high voltage generator described later. X-rays exposed from the X-ray focal point FP of the X-ray tube 21 are narrowed by the slit 23 and formed into a fan-shaped X-ray beam. The X-ray beam is then transmitted through the jaw JW of the subject P and attenuated, and the transmitted X-ray beam reflecting the attenuated state enters the detector 22.

  At the time of imaging, as shown in FIG. 2, the jaw JW of the subject P is positioned at a predetermined position in a three-dimensional imaging space IS defined between the X-ray tube 21 and the detector 22. For this reason, the X-ray tube 21 and the detector 22 face each other (face to face) across the jaw. The irradiated X-ray beam passes through the slit 23, then passes through the jaw portion JW (dentition, etc.) and is detected by the detector 22. Since the two arms 15 and 16 are rotationally driven by the rotating mechanism 13D at the time of imaging, the X-ray tube 21 and the detector 22 around the jaw part, respectively, around the jaw part along a predetermined circular orbit. Rotate. During the rotation, irradiation and detection of the X-ray beam are executed at predetermined intervals.

  When viewed along the X-axis direction facing the YZ plane, the X-ray tube 21 and the detector 22 are respectively driven to rotate along circular trajectories Tx and Td centered on the rotation center O determined in advance on the system side. The The radii Dx and Dd from the rotation center O to the circular trajectories Tx and Td are set to different values in consideration of X-ray exposure, detection accuracy, downsizing of the apparatus, mechanical interference with the patient, and the like. (See FIG. 2). In the present embodiment, Dx ≠ Dd, and particularly Dx> Dd. The reason why the distance (radius Dd) from the rotation center O to the detector 22 is smaller than that (radius Dx) from the rotation center O to the X-ray tube 21 is that the position of the detector 22 is set as much as possible. This is to reduce the attenuation of the incident intensity of X-rays. The distance (radius Dx) from the rotation center O to the X-ray tube 21 is set to a value that can ensure the distance between the X-ray tube and the skin defined by the standard.

  Therefore, the X-ray tube 21 and the detector 22 are always opposed to each other (facing to each other), and irradiation and detection of X-rays along a plurality of predetermined desired X-ray paths for the jaw JW (dentition) are performed. In order to perform the above, the X-ray tube 21 and the detector 22 are independently driven at different angular velocities.

  Note that “opposing each other” described above is a cone-like shape that is irradiated from the dotted X-ray focal point FP of the X-ray tube 21 when viewed along the X-axis direction as shown in FIG. The state in which the irradiation range of the shaped X-ray beam and the X-ray detection surface 22A (described later) of the detector 22 coincide with each other. In particular, the center line in the direction along the YZ plane of the X-ray beam includes an axis T that intersects the center position C in the width direction of the X-ray detection plane (the width in the direction along the YZ plane) at 90 °. The state is referred to as a “facing state” (see FIG. 3). In FIG. 3, a linear position extending in the Z-axis direction from the mechanical rotation center O is defined as a rotation angle θ = 0, and ± rotation directions are set clockwise and counterclockwise from this rotation position.

  For this reason, in order to realize the above-mentioned “always opposite (or facing each other)”, the opposing arm portions 15A and 16A including the X-ray tube 21 and the detector 22 out of the arms 15 and 16 have the axis Rotation (spinning, that is, posture) can be independently performed around AXs and AXd (see FIGS. 1 to 3). For this purpose, rotation driving mechanisms 15B and 16B such as motors are provided on the arms 15 and 16, respectively. The drive control of the rotation drive mechanisms 15B and 16B is executed by a controller of the console 17 described later.

  In the present embodiment, the positions of the intersection position C and the axis AXd are made to coincide with each other in the Y-axis direction. Further, the circular trajectories Tx and Td shown in FIG. 3 follow the positions of the axes AXs and AXd, respectively, when viewed in the YZ plane.

  As shown in FIG. 4, the detector 22 has an array (sensor circuit) of a plurality of detection modules B <b> 1 to Bm in which X-ray imaging elements are two-dimensionally arranged. The plurality of detection modules B1 to Bm are created as blocks independent from each other, and are mounted in a predetermined rectangular shape on a substrate (not shown) to form the entire detector 22.

A plurality of detection modules B1 to Bm are arranged in the vertical (X axis) direction (17 in the vertical direction) while providing a constant gap between the individual modules, and the individual modules are arranged in the scanning direction O _Y. The angle θ is set to, for example, about 14 °, and the ratio of the vertical and horizontal lengths created by the plurality of detection modules B1 to Bm is large. The elongated rectangular surface forms the X-ray detection surface 22 A. Since the detection modules B1 to Bm are arranged obliquely, the X-ray detection surface 22A is located inside the individual detection surfaces of the plurality of modules B1 to Bm. Of course, the angle θ may be set to 0 °.

  The structure of the detector 22 having this obliquely arranged detection module and the processing of the detection signal by the subpixel method are known, for example, from International Patent Publication WO 2012/0866648 A1.

  4 is a central axis when the detector 22 itself rotates (rotates).

  Each detection module B1 (~ Bm) is made of a semiconductor material that converts X-rays directly into electrical pulse signals. For this reason, the detector 22 is a photon counting X-ray detector of a direct conversion method using a semiconductor.

As described above, the detector 22 is formed as an array of a plurality of detection modules B1 to Bm. Each detection module Bm includes a detection circuit Cp (see FIG. 5) for detecting X-rays and a data counting circuit 51 _n (see FIG. 5) stacked together with the detection circuit Cp, as is well known. The detection circuit Cp includes, for each detection module, a semiconductor layer that directly converts X-rays into an electrical signal, and a charging electrode and a collecting electrode that are respectively stacked on both sides (not shown). X-rays are incident on the charged electrode. The charged electrode is a common electrode, and a high bias voltage is applied between the charged electrodes. The semiconductor layer and the collecting electrode are divided into a grid pattern, and by this division, a plurality of small regions are formed that are arranged in a two-dimensional array at a certain distance from each other. As a result, a plurality of stacked bodies of semiconductor cells C (see FIGS. 4 and 5) and collecting electrodes arranged in a two-dimensional manner on the charged electrode are formed. The plurality of stacked bodies to form a plurality of pixels S _n arranged in a two dimensional grid pattern.

As a result, a plurality of pixels S _n (n = 1 to N) occupying a predetermined area necessary for the detector 22 are formed by the entirety of the plurality of detection modules B1 to Bm. The plurality of pixels _{S n} constitutes a pixel group Cp (refer to FIG. 5).

The number of pixels detecting module B1~Bm each is 40 × 40 pixels, the size of each pixel _{S n} is 200 [mu] m × 200 [mu] m, for example. This pixel size is set to a value that allows detection of incident X-rays as a collection of many photons. Each pixel S _n is responsive to incident of each photon of X-ray, and outputs an electrical pulse of amplitude corresponding to the energy possessed by the photon. That is, each pixel S _n may convert the X-rays incident on that pixel directly, into electric signals.

Therefore, the detector 22, the photon constituting the cone beam-like X-rays incident, counts for each pixel S _n which constitute the detection surface of the detector 22, the quantity of electricity that reflects the count value For example, data is output at a high frame rate of 300 fps. This data is also called frame data.

As the semiconductor material of the semiconductor layer, that is, the semiconductor cell C, cadmium telluride semiconductor (CdTe semiconductor), cadmium zinc telluride semiconductor (CdZnTe semiconductor (CZT semiconductor)), silicon semiconductor (Si semiconductor), thallium bromide (T1Br) Mercury iodide (HgI ₂ ) or the like is used. Instead of this semiconductor cell, it is composed of a cell that combines a scintillator material that is subdivided into columns and optically shielded from each column, and a photoelectric converter composed of a combination of fine avalanche photodiodes. May be.

For this reason, when X-rays enter the semiconductor cell C, charges (electrons, holes) are generated inside the cell, and a pulse current corresponding to the amount of the charge flows. This pulse current is detected by the current collecting electrode. As a result, the amount of charge varies depending on the energy value of the X-ray photons. Therefore, the detector 22 outputs an electrical pulse signal corresponding to the energy value of the photons for respective pixels S _n.

The detector 22 further comprises respective semiconductor cell C, that the data counting circuit ₅₁ on the output side of each of the plurality of pixels _{S n} n a (n = 1~N). Here, each pixel _{S n,} i.e., a route to each of the data counting circuit ₅₁ 1 from each (to 51 _N) of the semiconductor cell C, and optionally, referred to as acquisition channels _CN n (n = _1~N) (See FIG. 5).

  The structure of this group of semiconductor cells C is also known from Japanese Patent Application Laid-Open Nos. 2000-69369, 2004-325183, and 2006-101926.

Incidentally, the size of each pixel S _n described above (200 [mu] m × 200 [mu] m) is adapted to a sufficiently small value that is capable of detecting X-rays as photons (particles). In the present embodiment, the size capable of detecting X-rays as the particles is “between electric pulses responding to each incident when a plurality of radiation (for example, X-ray) particles are successively incident at or near the same position. The occurrence of the superposition phenomenon (also called pile-up) is defined as “a size that can be substantially ignored or whose amount is predictable”.

  However, even with such a pixel size, it is not possible to avoid all occurrences of the superposition phenomenon. Even when two or more electrical pulses are both observed in the same pixel, the superposition phenomenon does not occur if they are separated from each other in time. On the other hand, when two or more electric pulses are difficult to separate in time in the same pixel, a superposition phenomenon occurs, and the two electric pulses overlap to be observed as one electric pulse having a high peak value. Is done.

When this superposition phenomenon occurs, X-ray particle countdown (also called pile-up count loss) occurs in the characteristic of “number of incidents versus actual count value” of X-ray particles. Therefore, the size of the pixel S _n to form the X-ray detector 12, the magnitude of which can be regarded as the counting loss does not occur or does not substantially occur, or are set to an extent counting the drop amount can be estimated .

Subsequently, a circuit electrically connected to the detector 22 will be described with reference to FIG. Each of the plurality of data counting circuits 51 _n (n = 1 to N) includes a charge amplifier 52 that receives an electrical signal of an analog amount output from each semiconductor cell C, and the waveform shaping is performed at the subsequent stage of the charge amplifier 52. Circuit 53, multi-stage comparator 54 _i (here i = 1 to 4), multi-stage counter 56 _i (here i = 1 to 4), multi-stage D / A converter 57 _i (here i = 1 to 4) 4) A latch circuit 58 and a serial converter 59 are provided.

Each charge amplifier 52 is connected to each current collecting electrode of each semiconductor cell S, charges up the current collected in response to the incidence of X-ray particles, and outputs it as a pulse signal of electric quantity. The output terminal of the charge amplifier 52 is connected to a waveform shaping circuit 53 whose gain and offset can be adjusted. The waveform of the detected pulse signal is processed with the previously adjusted gain and offset to shape the waveform. The gain and offset of the waveform shaping circuit 53, in consideration of the variation in non-uniformity and the circuit characteristics for charge-charge characteristic for each pixel S _n of semiconductor cell C, is calibrated. As a result, it is possible to increase the output of the waveform shaping signal from which non-uniformity has been eliminated, and the relative threshold setting accuracy. As a result, corresponding to each pixel S _n, i.e., the characteristics reflecting the energy value of the X-ray particle pulse signal waveform formatted output from the waveform shaping circuit 53 for each collection channel CN _n is substantially incident Have. Therefore, the variation between the collection channels CN _n is greatly improved.

The output terminal of the waveform shaping circuit 53 is connected to the comparison input terminals of the plurality of comparators 54 _{1 to} 54 ₄ . Analog reference thresholds (voltage values) th _i (here, i = 1 to 4) having different values are applied to the reference input terminals of the plurality of comparators 54 _{1 to} 54 ₄ as shown in FIG. Yes. This makes it possible to compare one pulse signal with each of the different analog amount thresholds th _{1 to} th ₄ . FIG. 6 shows the magnitude relationship (th ₁ <th ₂ <th ₃ <threshold) between the peak value (representing energy) of the pulse voltage generated in response to the input of one X-ray photon and the threshold values th _{1 to} th _4. th ₄ ) schematically.

The reason for this comparison is to examine which region (discrimination) the energy value of the incident X-ray particle enters among the energy regions set in advance divided into a plurality. It is determined which value of the analog quantity threshold th _{1 to} th ₄ exceeds the peak value of the pulse signal (that is, the energy value of the incident X-ray photon). Thereby, the energy area | region discriminated differs. Note that the lowest analog amount threshold th ₁ is normally set so as not to detect disturbances, noise caused by circuits such as the semiconductor cell S and the charge amplifier 42, or low-energy radiation that is not necessary for imaging. Is set as the threshold value. Further, the number of thresholds, i.e., the number of comparators is not necessarily limited to four, three, including the amount of the analog amount threshold th _1, or may be five or more.

Analog amount threshold th ₁ to TH ₄ described above, specifically, given from the calibration computing unit 38 of the console 17 for each pixel S _n in a digital value through the interface 32, i.e., for each acquisition channels. For this reason, the reference input ends of the comparators 54 _{1 to} 54 ₄ are connected to the output ends of the _four D / A converters 57 _{1 to} 574, respectively. The D / A converter ₅₇ 1-57 ₄ is connected to a threshold receiving end _T 1 via the latch circuit 58 (~T _N), the threshold receiving end _T 1 (~T _N) is the interface 32 of the console 17 It is connected.

The latch circuit 58 latches the threshold values th ₁ ′ to th ₄ ′ of digital quantities given from the threshold value applicator 40 via the interface 31 and the threshold value receiving end T ₁ (to T _N ) at the time of imaging, and corresponding D / are output to a converters ₅₇ 1 to 57 _4. Thus, D / A converter ₅₇ 1-57 ₄ may be provided to each comparator ₅₄ 1-54 ₄ threshold _th 1 to TH ₄ analog amounts commanded as a voltage amount. Each acquisition channel CN _n is one from the D / A converter 57 _i (i = 1 to 4) to the counter 56 _i (i = 1 to 4) via the comparator 54 _i (i = 1 to 4). Or it is connected to a plurality of circuit systems. This circuit system is called “discrimination circuit” DS _i (i = 1 to 4).

FIG. 7 shows a setting example of the energy threshold TH _i (i = 1 to 4) corresponding to the analog amount threshold th _i (i = 1 to 4). This energy threshold TH _i (i = 1 to 4) is, of course, a discriminating value that is set discretely and can be set to an arbitrary value by the user. FIG. 7 schematically shows an X-ray spectrum when an appropriate material is used for the anode material of the X-ray tube 21. The horizontal axis indicates the X-ray energy depending on the tube voltage of the X-ray tube 21, and the vertical axis indicates the incidence frequency of the X-ray photons. This incidence frequency is a factor representative of the count value (count) or intensity of X-ray photons.

The analog amount threshold th _i is an analog voltage applied to the comparator 54 _i in each discrimination circuit DS _i , and the energy threshold TH _i is an analog value for discriminating the X-ray energy (keV) of the energy spectrum. For the continuous spectrum shown in FIG. 7, the first analog amount threshold th ₁ is lower than the region where counting the number of X-ray photons is unnecessary (the region where there is no meaningful X-ray information for counting and circuit noise is mixed). The first energy region ER ₁ of the eye is set corresponding to an energy threshold TH ₁ that can be distinguished. Further, the second and third analog amount threshold values th ₂ and th ₃ are set so as to sequentially provide the second and third energy threshold values TH ₂ and TH ₃ which are higher than the first energy threshold value TH ₁ . . Further, the fourth energy threshold TH ₄ is set to an energy value equal to the applied voltage to the X-ray tube, in which X photon count value = 0 if there is no superposition phenomenon in the energy spectrum. Here, the fourth energy threshold TH _4, each pixel S _n, it is one important feature that is matched to the energy value as a count value = 0.

As a result, appropriate discrimination points based on the characteristics and design values of the energy spectrum are defined, and the energy regions ER _{1 to} ER ₄ are set.

These energy thresholds TH _i are determined so that one or more subjects as a reference are assumed and the count value for a predetermined time for each energy region is substantially constant.

Therefore, the output of the comparator ₅₄ 1-54 _3, as shown in FIG. 5 are respectively connected to a plurality of counters ₅₆ 1 to 56 ₄ of the input terminal.

Each of the counters ₅₆ 1 to 56 ₄ counts up every time the output of the comparator ₅₄ 1-54 ₃ (pulse) is turned on. Thus, the counters ₅₆ 1 (to 56 ₄₎ the accumulated value _W 1 of the X-ray photon number every fixed time with the energy of more than the energy value which is discriminated in the energy region _ER 1 (to Er ₄₎ in charge of the '( to _W-4 ') as it can be counted for each pixel _{S n.}

Specifically, this counting operation is determined by the relationship between the detection voltage V _dec (photon detection energy value) input to the _four comparators 54 _{1 to} 54 ₄ and the threshold values th _{1 to} th ₄ . That is, when the detection voltage _V dec _<th 1 _{~th 4} is a all comparators ₅₄ 1 to 54 ₄ output = off. In other words, the output = 0 of the pixel _{S n.} As a result, noise components smaller than the energy threshold TH ₁ defined as the input energy counting limit are not counted. This noise component corresponds to an energy value signal belonging to the non-countable region ERx in FIG.

However, if the detection voltage V _dec exceeds the minimum threshold th ₁ (V _dec ≧ th ₁ ), the number of photons is counted. If their relationship is _{V dec} ≧ th _1, the outputs of all the comparators ₅₄ 1 to 54 ₄ is turned on. That is, all of the counters ₅₆ 1 to 56 ₄ of the count value _{W 1'~W 4} 'is incremented.

If the relationship of V _dec ≧ th _2, the output of the second and subsequent stages of the three comparators ₅₄ 2-54 ₄ is turned on. Thus, three counters ₅₆ 2-56 ₄ count value _{W 2'~W 4} 'is incremented. If the relationship of V _dec ≧ th ₃ is established, the outputs of the third-stage and fourth-stage comparators 54 ₃ and 54 ₄ are turned on. As a result, the count values W ₃ ′ and W ₄ ′ of the two counters 56 ₃ and 56 ₄ are counted up.

_Furthermore, if the relationship between _{V dec} ≧ th _4, the output of the comparator module 54 ₄ of the fourth stage is turned on, only the counter 56 ₄ count value _{W 4} 'of the fourth stage is counted up. In this case, the energy value of the photon related to the input is a noise component belonging to the region ER ₄ exceeding the third high energy region ER ₃ , disturbance, etc., which is not suitable for imaging or counting. On the other hand, the count value W ₄ ′ can be used as information for estimating or excluding photons that have caused a superposition phenomenon or simultaneously incident photons.

As described above, in the present embodiment, the counters 56 _{1 to} 56 ₄ count the number of photons having energy exceeding the energy region ER ₁ (to ER ₄ ) to be counted by the counters 56 _{1 to} 564. For this reason, the number of X-ray photons having energy belonging to each of the _first to fourth energy regions ER _{1 to} ER ₄ , that is, the number of X-ray photons to be obtained for each energy region is expressed as W ₁ , W ₂ , W ₃ , W ₄ , the relationship between the count values W ₁ ′, W ₂ ′, W ₃ ′, and W ₄ ′ of the counters 56 _{1 to} 56 ₄ is
W ₁ = W ₁ '-W ₂ '
_{W 2 = W 2 '-W} 3'
W ₃ = W ₃ '-W ₄ '
It becomes. Note that W ₄ = W ₄ ′ is not calculated because it is meaningless information (that is, the energy region of the X-ray photon cannot be specified) due to the superposition phenomenon.

Therefore, the count values W _{1 to} W _{4 that} are truly desired are obtained by subtraction processing based on the above equation by a data processor described later. Ideally, W ₄ = W ₄ ′ = 0.

Thus, in the present embodiment, X-ray photon number _W 1 to _W-4 belonging to the first to fourth energy regions _ER 1 to Er ₄ respectively, the actual count value _{W 1'~W 4} ' Is obtained by calculation (subtraction). Therefore, output on the comparator 54 _{1-54 4,} a combination of off now events, namely circuit to decipher whether the incident X-ray photons belongs to which energy region ER1~ER4 becomes unnecessary. This simplifies the circuit configuration mounted on the data counting circuit 51 _n of the detector 22.

  In addition, the meaning of “collection” for each energy region of the number of X-ray photons according to the present application is the meaning of “obtaining by calculation” from the actual count value as described above, and for each energy region as in a modification example described later. Both meanings of directly “counting” the number of X-ray photons are included.

The counter 56 _{1-56 4} described above start and stop signals is supplied via a start-stop terminal T2 from below to the controller of the console 17. Counting for a fixed time is managed from the outside using a reset circuit included in the counter itself.

Thus, during a certain period of time until reset by a plurality of counters 56 ₁ to 56 _4, the number of photons of X-rays incident on the detector 22 is counted for each pixel S _n. The number of photons counted value _W k of the X-ray '(k = 1~4), after being output from the respective counters ₅₆ 1 to 56 ₄ in parallel as the count value of the digital quantity, serial format by serial converter 59 Is converted to The serial converter 59 ₁ is connected to the serial and the remaining serial converter ₅₉ 2-59 all acquisition channels _N. Therefore, the count of all digital content is output from the last channel of the serial converter 59 _N serially sent to the console 17 via the transmitting end T3.

  In the console 17, the interface 31 receives these count values and stores them in a storage unit described later.

In the present embodiment, it is integrally constructed in CMOS by the semiconductor cell C and the data counting circuit 51 _n corresponding to N pixels S _n described above ASIC (Application Specific Integrated Circuit). Of course, the data counting circuit 51 _n may be configured as a circuit or device separate from the group of semiconductor cells C.

  In the above embodiment, the plurality of detection modules B1 to Bm receive a light incident from the scintillator, which is optically connected to the scintillator array in which a plurality of scintillators processed into columnar shapes are bundled, and the scintillator array. A plurality of avalanche photodiodes are mounted on the light receiving surface, and the avalanche photodiodes belonging to the region are electrically connected by a quenching element for each rectangular region having a predetermined size corresponding to the cell on the light receiving surface. And a silicon photomultiplier.

  The material of the scintillator is LFS (lutetium silicate), GAGG: Ce (gadolinium aluminum gallium garnet), Pr: LuAG (praseodymium-added lutetium aluminum garnet), or the same decay time and light emission amount as the Pr: LuAG. It may be a material having a specific gravity.

  As shown in FIG. 8, the console 17 includes an interface (I / F) 31 that performs input and output of signals, a controller 33 that is communicably connected to the interface 31 via a bus 32, and a first storage unit 34, a data processor 35, a display device 36, an input device 37, a calibration calculator 38, a second storage unit 39, first to fourth ROMs 40A to 40D, and a threshold value assigner 41.

  The controller 33 controls the driving of the panoramic imaging device 1 in accordance with a program given in advance to the first ROM 40A. For this control, the command value is sent to the high voltage generator 42 that supplies a high voltage to the X-ray tube 21, the command value is sent to the opening drive unit 23D to change the opening area of the slit 23, and the calibration is performed. This includes a drive command to the computing unit 38 and control related to obstacle shadow removal described later. The first storage unit 34 stores frame data and image data that are count values sent from the detector 22 via the interface 31.

  The data processor 35 operates based on a program given in advance to the second ROM 40B under the control of the controller 33. This operation includes a process for removing the obstacle shadow, which will be described later. During panoramic photography, the data processor 35 uses the operation to add to the frame data stored in the first storage unit 34 tomosynthesis based on a known calculation method called shift and add. Implement the law. Thereby, the panoramic image of the tomographic plane with the oral cavity of the subject P is obtained. The display unit 36 is responsible for displaying an image to be created, information indicating the operation status of the apparatus, and operator operation information given via the input unit 37. The input device 37 is used by an operator to give information necessary for imaging to the apparatus.

Further, the calibration computing unit 38, under the control of the controller 33, operating under program embedded in advance in the third ROM40C, giving for each energy discriminator circuit for each pixel S _n in the data counting circuit, Calibrate digital quantity threshold for X-ray energy discrimination.

  Under the control of the controller 33, the threshold value applicator 41 calls the threshold value of the digital quantity stored in the second storage unit 39 for each pixel and for each discrimination circuit at the time of imaging, and uses the threshold value as a command value as an interface. 31 to the detector 22. In order to execute this process, the threshold value assigner 41 executes a program stored in advance in the fourth ROM 40D.

  The controller 33, the data processor 35, the calibration calculator 38, and the threshold value assigner 41 are all provided with a CPU (central processing unit) that operates according to a given program. Those programs are stored in advance in each of the first to fourth ROMs 40A to 40D.

In this embodiment, as known from International Publication No. WO2011 / 142343 (International Application No. PCT / JP2011 / 060731), the structure of the imaging space IS is analyzed using a phantom, and the collection channel CN _{n of the} detector 22 is analyzed. Is calibrated. This calibration is executed at an appropriate timing such as before imaging or during maintenance inspection.

  Specifically, a phantom (not shown) having a marker positioned on a predetermined standard tomographic plane and capable of imaging known position information with X-rays is arranged in the imaging space IS of the panoramic imaging apparatus 1. . X-ray transmission data from the X-ray officer 21 is collected by the detector 22 to create a panoramic image. From the known position information of the marker and the marker position information on the panoramic image, the distance information between the X-ray tube 21 and the detector 22 and the height information of the X-ray tube with respect to the detector are calculated. From this calculation result and collected data, various parameters that define the positional relationship of the X-ray tube, the detector, and the tomographic plane, with the amount of change in the position of the line connecting the X-ray tube and the detector, are calculated. Thereby, parameters necessary for 3D image reconstruction are calibrated. For this reason, the projection direction can be expressed three-dimensionally by grasping the structure of the imaging space IS three-dimensionally. Therefore, as long as the panoramic image is in focus, an accurate panoramic image can be constructed with no distortion or little distortion in the three-dimensionally expressed image.

[How to remove obstacle shadows]
Next, the principle of the panoramic image obstacle shadow removal method executed in the present embodiment will be described.

The tomosynthesis image g (i) (i represents a position perpendicular to the X-ray beam) is an object f _i (i) (j = 1, 2,..., N) divided into n layers in the X-ray beam direction. And the sum of degradation functions h _i (i) (j = 1, 2,..., N) generated by the superposition operation at that position.

  Here, all functions are considered to be after logarithm. That is, the tomosynthesis image g (i) is expressed as follows. * Is a symbol indicating convolution.

g (i) = f ₁ (i) * h ₁ (i) + f ₂ (i) * h ₂ (i) +… + f _j (i) * h _j (i) +… + f _n (i) * h _n (i)
In the tomosynthesis process, a specific layer g _j (i) is imaged by making h _j (i) of a specific layer a delta function and making the others as uniform functions as possible.

Now, the model is simplified, and a three-layer model composed of two specific layers f ₁ (i) and f ₂ (i) and the other f _r (i) is considered as follows.

g (i) = f ₁ (i) * h ₁ (i) + f ₂ (i) * h ₂ (i) + f _r (i) * h _r (i)
Here, when focusing on the second layer, the following equation is established.

h ₂ (i) = δ (i)
g ₂ (i) = f ₁ (i) * h ₁ (i) + f ₂ (i) * δ (i) + f _r (i) * h _r (i)
= F ₁ (i) * h ₁ (i) + f ₂ (i) + f _r (i) * h _r (i)
… (1)
Further, when focusing on the first layer, the following formula is established.

h ₁ (i) = δ (i)
g ₁ (i) = f ₁ (i) * δ (i) + f ₂ (i) * h ₂ (i) + f _r (i) * h _r (i)
＝ f ₁ (i) + f ₂ (i) * h ₂ (i) + f _r (i) * h _r (i)
… (2)
Assuming that the component of f _r (i) * h _r (i) is very small, the following equation is derived from equations (1) and (2).

g ₂ (i) = f ₁ (i) * h ₁ (i) + f ₂ (i) (3)
g ₁ (i) = f ₁ (i) + f ₂ (i) * h ₂ (i) (4)
If g ₁ (i) is an image focused on the dentition and g ₂ (i) is an image focused on the cervical spine, then the true dentition image f ₁ (i) after removal of the obstacle shadow is
f ₁ (i) = (g ₁ (i) -g ₂ (i) * h ₂ (i)) / (1-h ₁ (i) * h ₂ (i)) (5)
As given.

The meaning of equation (5) is
(Image focused on dentition)-
(Image focused on cervical spine * blurring function at dentition position)
This means that an image of only the dentition can be obtained by dividing by the superposition integration of the blur functions corresponding to the positions.

Even if the components of f _r (i) * h _r (i) are not small, the above calculation is valid if the bias component is subtracted from g ₁ (i) and g ₂ (i) as certain parameter values. Therefore, it is possible to remove or reduce the obstacle shadow.

[Specific flow of removal of obstacle shadow]
Next, the removal of obstacle shadows executed in accordance with the above-described principle in the panorama imaging apparatus 1 will be specifically described. Here, it is assumed that the obstacle shadow is a shadow of the cervical vertebra reflected in the panoramic image of the dentition of the subject.

  First, the head of the subject P, that is, its jaw JW is positioned at the initial position of the imaging space IS between the arms 15 and 16 as shown in FIGS. Thereafter, a scan is instructed by the controller 33, the X-ray tube 21 and the detector 22 are rotated around the jaw JW, and frame data output by the detector 22 at a predetermined cycle is collected (FIG. 9, step S1). ). This frame data is temporarily stored in the first storage unit 34. This data collection is performed by a conventionally known method. For example, as shown in FIG. 10, in this data collection, the X-ray tube 21 and the detector 22 are rotated around the jaw portion JW so as to be focused on a reference tomographic plane (cross section) preset in the dentition TR. Executed. The code | symbol CS shown in FIG. 10 shows a cervical spine.

  Next, the data processor 35 reconstructs a panoramic image focused on the reference tomographic plane (see FIG. 11A) of the dentition TR using the frame data temporarily stored in the first storage unit 34. Are displayed (step S2). An example of this panoramic image is shown in FIG. The reconstructed panoramic image data is stored in the first storage unit 34.

  This reference tomographic plane is, for example, a horseshoe-shaped tomographic plane that is set to pass through substantially the center of a statistically standard size and shape, which is known from US Publication US2006 / 0203959A1 and Japanese Patent Application Laid-Open No. 2007-136163. Cross section). However, a panoramic image to which this obstacle shadow removal method can be applied is not necessarily limited to a panoramic image strictly along the reference tomographic plane. For example, as known from US Publication No. 2012 / 0230467A1, the panoramic image reconstruction method described in Japanese Patent Laid-Open No. 2007-136163 may be further improved. In other words, there are usually individual differences in the dentition and tooth size and size of each subject, and the dentition and the shape and position of the teeth are mostly deviated from the reference tomographic plane. Even with such dentitions and teeth, according to the panoramic image reconstruction method described in US Publication No. 2012 / 0230467A1, the actual position is more accurately focused and drawn, and the enlargement ratio between teeth It is possible to provide a panoramic image in which the difference is corrected or reduced. As described above, a panoramic image of a cross section that is entirely or locally deviated from the reference tomographic plane can also be an object of the obstacle shadow removal method of the present application. In other words, the panoramic image to be subjected to this removal method is a tomographic plane along the dentition of the subject, and the tomographic plane may be uneven.

  The reference tomographic plane SS is a similar horseshoe shape, but a plurality of types having different sizes are prepared, considering that the subject has a difference in the size of the jaws such as adults and children. It is desirable to select a fault plane of any size.

  In step S2, specifically, the image of the optimum focus on the reference tomographic plane is reconstructed by the tomosynthesis method, that is, shift-and-add processing. The second ROM stores gain curve information for optimally focusing a reference tomographic plane of a predetermined dentition TR. An example of this gain curve is shown in FIG. The gain curve is a curve that indicates an amount (differential value of the curve) by which each piece of strip-shaped frame data (frame image) is shifted (shifted) with respect to each other in the shift-and-add process.

  The shift amount in the gain curve corresponds to the blur amount when creating the blur function at the dentition position. Therefore, the data processor 35 uses the shape of the blur as a Gaussian function, matches its standard deviation to this shift amount, and convolves with the frame data as a shift variant blur function that changes for each angular position. Thereafter, the data processor 35 adds frame data to each other at each position (mutual addition of pixel values) to create a panoramic image of the reference tomographic plane along the column TR.

  Next, as described above, the data processor 35 uses the frame data temporarily stored in the first storage unit 34 to focus on the tomographic plane passing through the cervical vertebra CS (see FIG. 11A). Is reconstructed and displayed (step S3). An example of this panoramic image is shown in FIG. The reconstructed panoramic image data is also stored in the first storage unit 34.

The cervical vertebrae CS focused on the image creates a gain curve assuming a trajectory T _CS folded back trajectory T _TR focus on teeth surfaces. At this time, a gain curve is created depending on the position of the wire (lead phantom) of the calibration phantom. In this case, the gain curve is created in the form of inverting the above-described gain curve (see FIG. 11A). Based on the amount of deviation indicated by the gain curve (that is, the differential value of the curve), a shift-and-add operation is performed. In this case, generally, in the reconstruction of the dentition TR, the shift amount in the vicinity of the front teeth is small, but in the reconstruction of the cervical vertebra CS, the collected frame data is moved greatly to perform the addition operation.

The frame data used for reconstructing the panoramic image in steps S2 and S3 may be frame data having energy belonging to all energy regions ER _{1 to} ER ₃ , or may be included in a part of the energy region. It may be frame data of the energy to which it belongs. Even when using frame data of energy over all energy regions, in addition to simple averaging, weighting can be performed for each region.

  Since the detector 22 is a photon counting type, the frame data can be detected in a state where energy is discriminated for each pixel. This makes it possible to use transmission data (frame data) that sensitively reflects parameters such as a change in radiation quality when X-ray photons pass through the jaw material. For this reason, a panoramic image in which a specific substance in the jaw is emphasized can be acquired, and the obstacle shadow removal method can be performed based on the panoramic image.

Next, the data processor 35 performs a scaling process (step S4). Figure 13 shows the best focus image I _TR along the reference tomographic plane of the dentition TR with the vertical line. This vertical line shows the position of the specific angle measured with the calibration phantom. Further, FIG. 14 shows the best focus image I _CS along the fault plane cervical CS with the vertical line. This vertical line also indicates the position of a specific angle measured with the calibration phantom.

  In order to remove obstacle shadows such as the cervical spine, FIGS. 13 and 14 must be scaled to equal image sizes. The distance between the vertical lines in FIG. 13 and the distance between the vertical lines in FIG. 2 are different for each position, and the size of FIG. 14 cannot be matched to the size of FIG. 13 using one reduction ratio. Therefore, the scales of the two images are adjusted using the frame defined by the vertical lines. In order to easily match, it is only necessary to match each area surrounded by two vertical lines so as to have the same size. Further, in order to match with higher accuracy, it is desirable to match the corresponding positions for each strip-shaped frame data.

Here, the plurality of rectangular regions defined by vertical lines 15 represent components of the optimum focus image I _TR dentition TR, 15 present those regions by the vertical line of best focus image I _CS cervical CS Each area of the image _ICS is reduced so as to match. Thus, the size of both images _{I TR} and _{I CS} are the same.

Next, the data processor 35 executes a blurring process for the optimally focused image I _CS of the cervical vertebra CS (step S5). In order to determine the blur function in this operation, a gain curve is used. More specifically, the full width at half maximum of the gain value at each angular position of the gain curve: create a Gaussian function with a (FWHM full width at half maximum) , which is convolution optimally focused image I _CS cervical CS.

In the cervical vertebra blurred image I _CS ′ (not shown) thus formed, the dentition is greatly blurred, whereas the cervical vertebra is in contrast to the cervical vertebra in the image focused on the dentition surface. The blur is about the same.

  This blurring process, that is, the process of step S5 may be omitted depending on the situation.

After this blurring process, the data processor 35 performs, for each pixel, between the optimal focus image I _TR of the dentition _TR and the optimal focus image I _CS ′ of the cervical vertebra CS reduced and blurred as described above. The pixel value is subtracted or divided to perform a process of taking the image difference (step S6). When pixel values are expressed in natural logarithm, subtraction is performed, and “I _TR -I _CS ′” is performed between pixel values. On the other hand, when the natural logarithm of the pixel value is not taken, the division of “I _TR / I _CS ′” is performed between the pixel values. Further, since the density unevenness depending on the blur function occurs in the pixel value, the blur function on the cervical vertebra focal plane and the blur function on the dentition surface are superimposed and integrated, and a value obtained by subtracting this from 1 is used. By dividing each angular position, it is possible to obtain an image in which the influence of the cervical spine is reduced as much as possible and the dentition surface is focused. It is desirable to add this process to step S6.

Since the difference image by the difference or the division becomes the optimum focus image I _{TR_REV} of the dentition TR after removing (or reducing) the shadow of the cervical spine, this is displayed on the display 36 and, for example, the first image (Step S7).

FIG. 15 shows another example of the optimum focus image I _TR of the dentition TR before cervical vertebra removal. In this image _ITR , a white shadow (disturbance shadow) of the cervical vertebra CS is reflected in the central portion thereof, the vicinity of the front teeth is blurred, and the drawing ability is low. On the other hand, as shown in FIG. 16, in the optimally focused image I _{TR_REV} of the dentition TR after cervical vertebra removal, the cervical vertebra image is almost removed, and the anterior teeth are more clearly depicted accordingly.

As described above, according to this embodiment, a relatively simple correction after data acquisition, shading cervical CS that interfere shadow is reliably removed or alleviated from the panoramic image I _TR dentition TR. For this reason, the visualization ability of the dentition TR is improved and the contribution to the diagnosis is also great.

  In addition, the panoramic imaging apparatus 1 according to the present embodiment employs an imaging system in which the X-ray tube 21 and the detector 22 can rotate independently of each other so as to draw a circular orbit around the same center point O. Yes. For this reason, the advantage of being able to design the rotational trajectory of the X-ray tube 21 and the detector 22 while avoiding the cervical vertebra CS as an obstacle shadow as much as possible can be enjoyed as it is. The advantage of this trajectory design also doubles the effect of reducing obstacle shadows.

[Other embodiments]
In the above-described embodiment, the dentition TR is a structure for imaging, and the removal process is performed by using the reflection of the cervical vertebra CS in the panoramic image of the dentition TR as an obstacle shadow. However, this relationship may be reversed. . In other words, the cervical vertebra CS may be a structure for imaging, and the removal process may be performed using the reflection of the dentition TR in the panoramic image of the cervical vertebra CS as an obstacle shadow. In this case, the size of the panoramic image of the tomographic plane of the cervical vertebra CS is reduced to match the size of the panoramic image of the tomographic plane of the dentition TR. The point to apply is different. Other processing procedures are the same as or equivalent to those described above.

  Further, the obstacle shadow is not necessarily the cervical vertebra or the dentition, but may be, for example, the jawbone. For example, the dentition TR may be a structure to be imaged, and the removal process may be performed in the same manner as described above with the reflection of the left and right jaw bones in the panoramic image of the tomographic plane of the dentition TR as an obstacle shadow.

  In addition, the tomographic plane of the maxillary sinus of the jaw, the carotid artery, or the cortical bone dentition under the mandibular foramen is designated as the tomographic plane of one structure, and the cervical vertebra of the jaw is designated as the tomographic plane of the other structure. Alternatively, a tomographic plane passing through the hyoid bone may be used. Also in this case, the removal method similar to that described above can be implemented by capturing any one of the structures in the panoramic image of the other structure as an obstacle shadow.

  The above-described panoramic imaging apparatus 1 is an apparatus that captures an image in a state where the patient is lying on his / her back on the dental chair (in a supine position). However, the X-ray imaging apparatus according to the present invention has an imaging system that rotates with the jaw of the subject positioned between the X-ray tube and the detector, regardless of imaging in such a posture. For example, such an imaging system may be a device configuration in which such an imaging system is fixedly installed on the back surface of a chair or on a support column, and a patient receives an image while sitting or standing. Moreover, such an imaging system may be attached to a fixed structure such as a house or a vehicle wall or ceiling. Furthermore, such an imaging system may be configured as a portable unit, and may be configured to perform imaging by placing it on a patient's shoulder or installing behind a general chair.

  In the panoramic imaging apparatus 1 described above, the X-ray tube 21 and the detector 22 are provided with an imaging system that can be driven to rotate independently of each other while changing the distance between them around the same rotation center. The imaging system is not necessarily limited to such a configuration. As well known in the art, the obstacle shadow removal method of the present application can be applied even to an imaging system disclosed in, for example, Japanese Patent Application Laid-Open No. 2007-136163. In other words, the imaging system may rotate around the subject while the distance between the X-ray tube and the detector is always kept constant in a state where the X-ray tube and the detector face each other with the subject interposed therebetween.

  Furthermore, in the embodiment of the present application, the detector 22 is a photon counting type detector. However, the scintillator and the photoelectric element are combined to accumulate electrical signals for a predetermined time and output frame data. A so-called integral type detector may be used.

  Furthermore, the obstacle shadow removal method of the present application can also be applied to an X-ray imaging apparatus using a detector using a scintillator and a CCD (charge coupled device), to which the tomosynthesis method cannot be applied. In that case, an image taken under the rotational trajectory of the X-ray tube and detector focused on the cross-section passing through the dentition, and the rotational trajectory of the X-ray tube and detector focused on the cross-section passing through the cervical spine It is sufficient to prepare images taken under the image separately and perform the above-described obstacle shadow removal method between these images.

  The present invention is not limited to the configurations shown in the above-described embodiments and modifications, and can be implemented with various modifications without departing from the gist of the claims.

1 Dental panoramic imaging device 13D Rotating mechanism (supporting means)
17 Console
15,16 arm (support means)
21 X-ray tube 22 Detector 23 Slit 33 Controller (one of the elements for functionally realizing various means)
34 1st memory | storage part 35 Data processor (one of the elements which implement | achieves various means functionally)
36 Display 37 Input device 40A-40D ROM
51 _n , 151 _n Data counting circuit 54 Comparator 55 Energy domain allocating circuit 56 Counter 57 D / A converter 58 Latch circuit 59 Serial converter C Semiconductor cell Cp detection circuit S _n pixel DS _i discrimination circuit CN _n data collection channel

Claims (8)

An imaging system comprising an X-ray tube for irradiating X-rays and an X-ray detector for detecting the X-rays, with the jaw portion of the subject positioned between the X-ray tube and the detector In a dental X-ray imaging apparatus adapted to rotate,
First image acquisition means for acquiring a first panoramic image focused on one of the two structures existing on the jaw;
Second image acquisition means for acquiring a second panoramic image focused on the remaining one of the two structures;
Based on the first and second panoramic images, when the one structure is an object to be imaged and the remaining one structure is reflected on the panorama image of the object as an obstacle shadow. Object image generation means for generating a panoramic image of the object from which the reflection of the obstacle shadow is removed or reduced,
A dental X-ray imaging apparatus comprising: The first image acquisition means is configured to acquire a panoramic image of a dentition focused on the dentition of the jaw as the first panoramic image,
The second image acquisition means is configured to acquire a panoramic image of a jaw focused on a tomographic plane passing through the cervical vertebra of the jaw as the second panoramic image.
The dental X-ray imaging apparatus according to claim 1.   The first image acquisition means is a panoramic image of a tomographic plane based on a predetermined reference tomographic plane passing through the dentition or a deviation of the actual position of the dentition from the reference tomographic plane as the panoramic image of the dentition. The dental X-ray imaging apparatus according to claim 2, which is configured to acquire The object image generation means includes:
Scaling means for adjusting the size of the panoramic image of the cervical spine to the size of the panoramic image of the dentition;
A blur image generating unit that generates a blur image of the cervical spine by superimposing and integrating a blur function defined by the focal plane on each pixel of the panoramic image of the cervical spine that has been adjusted in size by the scaling unit;
The pixel value is subtracted or divided for each pixel between the panoramic image of the dentition acquired by the first image acquisition unit and the blurred image of the panoramic image of the cervical vertebra generated by the blur image generation unit. Difference image calculation means for providing the difference image as a panoramic image of the object after the obstacle shadow has been removed or reduced,
The dental X-ray imaging apparatus according to claim 2, wherein the dental X-ray imaging apparatus is provided.   5. The dental X-ray imaging apparatus according to claim 4, wherein the blur function is set to a smaller value in a portion where the cervical vertebra is depicted on the panoramic image of the cervical vertebra. . The first image acquisition means is configured to acquire a panoramic image of a dentition focused on a tomographic plane passing through the cervical vertebra of the jaw as the first panoramic image,
The second image acquisition means is configured to acquire a panoramic image of the jaw focused on the dentition of the jaw as the second panoramic image.
The dental X-ray imaging apparatus according to claim 1.   The detector includes a plurality of pixels for discriminating and collecting electric pulses corresponding to the energy of the photons for each incident of the photons of the X-rays, and arranging the pixels in a plane The dental X-ray imaging apparatus according to claim 1, wherein the dental X-ray imaging apparatus is a photon counting type detector having a detection surface. An imaging system comprising an X-ray tube for irradiating X-rays and an X-ray detector for detecting the X-rays, with the jaw portion of the subject positioned between the X-ray tube and the detector In an image correction method in dental X-ray imaging that is rotated,
Obtaining a first panoramic image focused on one of the two structures present in the jaw;
Obtaining a second panoramic image focusing on the remaining one of the two structures;
Based on the first and second panoramic images, when the one structure is an object to be imaged and reflection of the remaining one structure on the panoramic image of the object is an obstacle shadow Generating a panoramic image of the object with the obstruction shadows removed or reduced;
An image correction method in dental X-ray imaging characterized by the above.